Single photon emission computed tomography (SPECT) is a nuclear medicine tomographic imaging technique that has traditionally used homogeneous area radiation detectors to detect gamma ray emissions. Conventionally, this imaging technique accumulates counts of gamma photons that are absorbed by a scintillation crystal. The crystal scintillates in response to interaction with gamma radiation to produce a flash of light. Photomultiplier tubes (PMTs) behind the scintillation crystal detect the flashes of light and a computer sums the fluorescent counts. The sum of fluorescent counts is a measure of the energy of the detected gamma photon. The computer in turn constructs and displays an image of the relative spatial count density of detected gamma photons on a monitor. Images can be formed by detected gamma photons within user-specified energy limits. These images then reflect the distributions and relative concentrations of multiple radioactive tracer elements or multiple gamma photons with differing energies emitted from the same radioactive tracer element that is present in the organs and tissues imaged.
In more detail, U.S. Pat. No. 3,011,057 for RADIATION IMAGE DEVICE, by Hal O. Anger, which is incorporated by reference herein in its entirety, describes a nuclear medicine imaging device that uses a single sodium iodide (NaI) scintillation crystal to detect the gamma ray emissions. Here, the radiation imaging device generally includes a detector, such as a scintillation crystal, for transforming gamma ray emissions to light photons in response to incident gamma ray events, and a photodetector to detect the light photons emitted from the scintillation crystal. The photodetector, typically a photomultiplier tube that is optically coupled to the scintillation crystal, detects a fraction of the scintillation photons produced from absorption of a gamma ray into the scintillation crystal and produces an electronic current that is proportional to the number of detected scintillation photons.
In a technique used with a NaI detector, a nuclear medicine imaging device forms a high-resolution image through the use of a single-aperture collimator that provides collimated gamma ray paths to the detector. In this technique, the position of the gamma ray emission at the point of absorption in the scintillation crystal is determined by an algorithm based on the magnitude of electric signals from each of a plurality of photomultiplier tubes positioned over the crystal. This algorithm can be implemented by use of a resistor matrix connecting the outputs of the photomultiplier tubes. For close proximity images, a single long-bore, small-aperture collimator hole can be used, with the collimator being scanned over the radiation field of interest in a two-dimensional scanning manner, to thereby sample radiation distribution over each of the image points in the radiation field.
In addition to the above-described NaI detectors, nuclear medicine imaging devices with pixellated radiation detector elements, typically cadmium zinc telluride (“CZT”) crystals, have recently been developed. In more detail, U.S. Pat. No. 6,838,672 for HIGH RESOLUTION, MULTIPLE DETECTOR TOMOGRAPHIC RADIONUCLIDE IMAGING BASED UPON SEPARATED RADIATION DETECTION ELEMENTS, by Douglas J. Wagenaar et al., which is incorporated by reference herein in its entirety, describes a pixellated radiation detector. Here, the pixellated detector is generally characterized by multiple detector elements.
In the above described nuclear medicine imaging devices, the performance of the imaging devices can be improved through the use of multiple radiation detectors. Also, multiple holes can be used with a collimator to form a multi-hole collimator to further increase the number of counts obtained at each point, provided they are sufficiently separated from each other such that detected counts can be associated with a particular collimator hole. The use of multiple detectors and/or a multi-hole collimator is advantageous because the nuclear medicine imaging device may collect samples from a target in less time. An imager having two detectors and/or a two-hole collimator, for instance, may scan a target twice as fast as an imager having a single detector. Furthermore, the use of multiple detectors to scan a target may improve the resolution of the scanning by reducing the variance and resulting statistical error produced by a single detector. However, configuring the multiple detectors and/or multi-hole collimator for precisely scanning throughout a sampling area adds still greater complexity to the design for the nuclear medicine imaging device. As a result, there exists a need for a nuclear medicine imaging device having multiple detectors and/or a multi-hole collimator to have a relatively simple design.
Also, relatively more complicated iterative 3D image reconstruction techniques need to be used with nonstandard pinhole data acquisition geometries versus a standard pinhole SPECT imaging geometries involving the use of a pinhole collimator with a single pinhole aperture, while tomographic projection data are acquired on a planar circular orbit with a flat 2D detector. Examples of the nonstandard pinhole data acquisition geometries include the use of double or triple detectors, multi-pinhole collimator, helical scanning orbits, etc. That is, to implement image reconstruction techniques, there should be accurate geometric descriptions of the projection operators for these nonstandard pinhole data acquisition geometries, which can be very complex when considering geometric misalignments caused by mechanical imperfections. It is also crucial to estimate and then correct for geometric misalignments (a.k.a. geometric calibrations) in pinhole or cone-beam tomographic imaging geometries in order to minimize degradations and/or artifacts in the reconstructed images. Therefore, geometric calibrations and image reconstructions are closely knitted topics in pinhole SPECT imaging, and a reconstruction method implemented without the capability of correcting for geometric misalignments may be of little use in practical molecular imaging studies.
As such, to facilitate the translation of geometric calibration methods and improve the efficiency of implementing image reconstruction algorithms capable of correcting for geometric misalignments, there is a need to unify the geometric descriptions of projection operators for standard and nonstandard pinhole SPECT imaging geometries. Doing so may potentially speed up the implementation of effective reconstruction algorithms and the investigation of accurate geometric calibration methods for a nonstandard geometry.
The above information disclosed in this Background section is only for enhancement of understanding of the background of the invention and therefore it may contain information that does not form the prior art that is already known in this country to a person of ordinary skill in the art.